Apparatus and method for monitoring pressure related changes in the extra-thoracic arterial circulatory system

ABSTRACT

A method and apparatus for monitoring changes in the intra-thoracic pressure of a patient due to the patient&#39;s respiratory activity or volumetric changes in the extra-thoracic arterial circulatory system due to cardiac function based on the changes in pressure in the patient&#39;s extra-thoracic arterial circulatory system as measured by a plethysmography sensor, such as an photoplethysmograph. A frequency spectrum is generated for the plethysmograph signal and the frequencies of interest is isolated from the frequency spectrum by setting appropriate cutoff frequencies for the frequency spectrum. This isolated frequency is used to filter the plethysmograph signal to provide a signal indicative of the patient&#39;s respiratory activity or cardiac function.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Divisional under 35 U.S.C. §121 of U.S. patentapplication Ser. No. 10/999,186, filed 29 Nov. 2004, which claimspriority under 35 U.S.C. §119(e) from provisional U.S. patentapplication No. 60/525,954 filed Dec. 1, 2003, the contents of each ofwhich are incorporated herein by reference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention pertains to a method and apparatus for monitoringchanges in the intra-thoracic pressure of a patient due to the patient'srespiratory activity or cardiac function, and, in particular, to a firsttechnique in which pleural pressure changes due to respiratory effortare monitored based on the changes in pressure in the patient'sextra-thoracic arterial circulatory system, and to a second technique inwhich a patient's vessel distention in the extra-thoracic arterialcirculatory system due to respiratory activity or cardiac function aremonitored.

2. Description of the Related Art

Numerous patients arrive at a hospital's emergency room each daycomplaining of a respiratory disorder, such as difficulty breathing,wheezing, shortness of breath, etc. Many of these patient's areincapable of communicating effectively with their physician, for examplethey may be too young, incapacitated in some way, or have a mentaldeficiency that prevents effective communication with their caregivers.It would be desirable in such situations if a technique existed formonitoring their respiratory function independent of the patient'sability to communicate, i.e., with regard to the patient's descriptionof the problem. Such a technique would also serve as an objectiveevaluator of a patient's condition, even if subjective communicationwere possible.

Conventional methods of assessing respiratory function, including workof breathing, include visually monitoring the respiratory effort of thepatient, for example, by observing whether the patient is havingdifficulty breathing. This provides no objective, measurable indicationof the patient's well-being.

A more invasive, yet more objective pulmonary effort measuring techniqueinvolves placing an esophageal catheter in the patient's airway andmonitoring the pressure within the patient's esophagus. It is alsopossible to monitor a patient's work of breathing using a mechanicalventilator. However, this requires attaching the patient to theventilator. These methods are invasive and, therefore, have limitedapplication. For example, when an asthma patient enters the emergencydepartment of a hospital, he or she is usually not on a ventilator, yetwork of breathing needs to be assessed and treated immediately. In theICU, a significant number of patients are at high risk for respiratoryfailure or have recently been extubated. These patients are not on aventilator, yet monitoring their work of breathing weighs significantlyin the plan of care prescribed for them.

There is also a tremendous need to understand interactions between theheart and lungs of patients in the ICU. For example, any obstructive orrestrictive disease, such as chronic obstructive pulmonary disease(COPD) or congestive heart failure (CHF), will result in increasedintra-thoracic pressure swings. If the patient's work of breathing ishigh, blood flow from the heart changes within each breath. To date, atool does not exist that can illustrate these interactions. Anotherexample occurs when high ventilator pressures are needed. With eachventilator breath, blood flow from the heart changes within each breath.Thus, it is important to determine how low the ventilator pressures needto be to provide adequate ventilation without altering blood flow fromthe heart. This determination is very difficult to make because thedetermination will be different for each patient. Without an objectivemeasurement of the hemodynamic effect, this determination cannot bemade.

Finally, it is known to monitor the blood pressure of a patient todetect a symptom of a heart disease. For example, it is known to monitora patient's blood pressure for pulsus paradoxis, which is a greater thannormal decrease in systolic pressure and pulse wave amplitude duringinspiration. Pulsus paradoxis is associated with circumstances in whichrespiration is labored and often accompanies such conditions asemphysema, pulmonary embolus, cardiac tamponade, lung cancer, or CHF.Other symptoms of heart disease include:

-   -   (1) “waterhammer” pulse, which is associated with aortic        insufficiency, and is characterized by a rapid pressure upstroke        and rapid fall into diastole;    -   (2) anacrotic pulse, which is associated with aortic stenosis        and characterized by a delayed pulse upstroke;    -   (3) dicrotic pulse, which is associated with decreased arterial        tone and is characterized by an accentuated secondary pulse wave        that may feel like heart rate is twice as fast as normal;    -   (4) pulsus bisferiens, which is associated with combined aortic        stenosis and insufficiency and is characterized by double peaks        in the pulse waveform; and    -   (5) pulsus alternans, which is usually associated with heart        failure and is characterized by a large pulse wave followed by a        small secondary wave.

Conventional non-invasive blood pressure monitors are only capable oftaking a “snap shot” of the patient's blood pressure, i.e., the peaksystole and diastole pressure, each time the blood pressure is measured.Thus, they are not suited to detect the dynamic blood pressure changesassociated with these blood pressure related symptoms of heart disease.

It is known to monitor the blood pressure continuously, so that bloodpressure related symptoms of heart disease, such as pulsus paradoxis,can be readily detected. However, conventional continuous blood pressuremonitors are invasive; requiring locating a pressure sensor within thepatient's arterial circulatory system.

SUMMARY OF THE INVENTION

Accordingly, it is an object of the present invention to provide acardiopulmonary monitoring system that overcomes the shortcomings ofconventional monitoring techniques. This object is achieved according toone embodiment of the present invention by providing an extra-thoracicmonitoring system that includes a sensing means to detect aphysiological characteristic of a patient associated with pressurechanges in such a patient's circulatory system and for outputting afirst signal indicative of such pressure changes. The system alsoincludes a processing means to produce a thoracic pressure signal as ameasure of such a patient's intra-thoracic pressure due to respiration.This is accomplished in the processing means by isolating cardiacrelated pressure variations in the first signal from the sensing means.

It is yet another object of the present invention to provide a method ofmonitoring the pulmonary function of a patient that does not suffer fromthe disadvantages associated with conventional monitoring techniques.This object is achieved by providing a method that includes detecting aphysiological characteristic of a patient associated with pressurechanges in such a patient's circulatory system and for outputting afirst signal indicative of such pressure changes. The method alsoincludes producing a thoracic pressure signal as a measure of such apatient's thoracic pressure due to respiration by isolating cardiacrelated pressure variations in the first signal.

It is a further object of the present invention to provide a system andmethod for measuring a fractional concentration of oxygen inhaled by apatient (FO₂). This measurement technique can be used alone or inconjunction with the cardiopulmonary monitoring system discussed above.The FO₂ monitoring system includes a patient circuit adapted tocommunicate a flow of breathing gas to an airway of a patient and afirst flow sensor associated with the patient circuit. The first flowsensor quantitatively measures a flow of gas (Q_(T)) inhaled, exhaled,or inhaled and exhaled by a patient. The FO₂ monitoring system alsoincludes an oxygen conduit adapted to be coupled to an oxygen source andto the patient circuit to communicate oxygen from the oxygen source tosuch a patient. A second flow sensor is associated with the oxygenconduit to quantitatively measure a flow of the oxygen (Q_(O2)) in theoxygen conduit. A processing system determines the FO₂ based on theoutput of the first flow sensor and the second flow sensor. Theprocessing system is also capable of determining the average fractionalconcentration of oxygen inhaled by a patient over one breath (FIO₂) byidentifying the respiratory cycle.

It is yet another object of the present invention to provide a systemand method for displaying patient information, which can be used aloneor in combination with the pulmonary monitoring system or method and/orthe FO₂/FIO₂ monitoring system and method discussed above. The patientinformation display system includes means for determining fractionalconcentration of oxygen inhaled by a patient during one breathing cycle,means for measuring a pulse oximetry arterial oxygen saturation (SpO₂)of such a patient, a display having a display area, and a displaycontroller. This display controller causes a nomogram illustrating arelationship between the measured SpO₂, the FIO₂, and an estimatedshunt, to be displayed in a first field on the display area. Thenomogram shows the SpO₂ on a first axis, the FIO₂ on a second axis, anda plurality of curves. Each curve corresponds to a common estimatedshunt percentage. The display controller causes an indicator to bedisplayed on the nomogram at a location defined by coordinatescorresponding to a current value of the SpO₂ and the FIO₂. This providesa readily visible indication of the estimated shunt based on themeasured SpO₂ and the FIO₂ values.

It is a still further object of the present invention to provide acardiac monitoring system that overcomes the disadvantages associatedwith conventional blood pressure monitoring techniques. This object isachieved according to one embodiment of the present invention byproviding a non-invasive cardiac monitoring system that includes aphotoemitter adapted to direct light into the tissue of a patient andthrough a portion of the patient and a photodetector adapted to receivelight after having been transmitted through or having been reflectedfrom a portion of such a patient. A processor produces a cardiacpressure signal as a measure of the patient's vascular pressure due tocardiac function by isolating cardiac related pressure variations in thefirst signal. This enables the present invention to monitor changes inthe patient's blood pressure non-invasively and substantiallycontinuously for detecting symptoms of cardiac dysfunction. The systemof the present invention is also capable of measuring heart rate andmonitoring heart rate variations that occur within each breath.

It is yet another object of the present invention to provide a method ofmonitoring the cardiac function of a patient that does not suffer fromthe disadvantages associated with conventional blood pressure monitoringtechniques. This object is achieved by providing a method that includes:(1) passing light through a portion of the patient, (2) receiving lightafter having been passed through the patient, (3) outputting a firstsignal based on the received light, and (4) producing a cardiac pressuresignal as a measure of the patient's vascular pressure due to cardiacfunction by isolating cardiac related pressure variations in the firstsignal. As noted above, this enables the present invention to monitorthe patient's blood pressure non-invasively on a substantiallycontinuous basis for detecting a symptom of a cardiac disorder.

These and other objects, features and characteristics of the presentinvention, as well as the methods of operation and functions of therelated elements of structure and the combination of parts and economiesof manufacture, will become more apparent upon consideration of thefollowing description and the appended claims with reference to theaccompanying drawings, all of which form a part of this specification,wherein like reference numerals designate corresponding parts in thevarious figures. It is to be expressly understood, however, that thedrawings are for the purpose of illustration and description only andare not intended as a definition of the limits of the invention. As usedin the specification and in the claims, the singular form of “a”, “an”,and “the” include plural referents unless the context clearly dictatesotherwise.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic diagram of the basic components of anextra-thoracic monitoring system according to the principles of thepresent invention;

FIG. 2 is a perspective view of the extra-thoracic monitoring systemshowing the housing and one embodiment of a sensor suitable for use withthis monitor;

FIG. 3 is a schematic diagram of a first embodiment of theextra-thoracic monitoring system according to the present invention;

FIG. 4 is a schematic diagram of the hardware components of theextra-thoracic monitoring system of FIG. 3;

FIG. 5 is a detailed circuit diagram corresponding to the schematicdiagram of FIG. 4;

FIGS. 6A-6K illustrate portions of the circuit shown in FIG. 5;

FIG. 7 is a timing diagram for the operation of the circuit shown inFIG. 5;

FIG. 8 is a schematic diagram of a user's tissue disposed between aphotoemitter and a photodetector of the present invention;

FIG. 9 is a schematic diagram illustrating an exemplary embodiment ofthe extra-thoracic monitoring system according to the principles of thepresent invention;

FIG. 10 is a graph of a raw hypothetical plethysmograph signal detectedby the plethysmograph monitoring portion of the extra-thoracicmonitoring system;

FIG. 11 is a graph of a respiratory component contained in the rawplethysmograph signal of FIG. 10;

FIGS. 12A and 12B are graphs illustrating an exemplary Fourier transformof the signal of FIG. 10;

FIGS. 13A and 13B are graphs illustrating an exemplary Fast FourierTransform (FFT) of the signal of FIG. 10 showing the selection of cutofffrequencies used in processing the plethysmograph signals according tothe principles of the present invention;

FIG. 14 is a graph illustrating a raw NIVD waveform, an NIVD_(Thoracic)waveform, and an NIVD_(Cardiac) waveform;

FIG. 15 is a graph showing a relationship between the raw NIVD signalpeak-to-peak values and breathing frequency;

FIGS. 16-19 are graphs showing an impact of a patient'sinspiratory-to-expiratory ratio on the FFT frequency spectrum signal;

FIG. 20 is a graph illustrating intra-breath pulse rate variations;

FIG. 21 is a graph illustrating a technique for determining the pulsusparadoxis of a patient using the extra-thoracic monitoring system of thepresent invention;

FIGS. 22 and 23 are schematic diagrams of two embodiments of systems formeasuring the fractional concentration of oxygen inhaled by a patientsuitable for use with the extra-thoracic monitoring system of thepresent invention;

FIG. 24 illustrates an exemplary nomograph display according to theprinciples of the present invention;

FIGS. 25-33 are screen shots of visual displays in a user interface foruse with the extra-thoracic monitoring system of the present invention;and

FIG. 34 is a graph illustrating the relationship between extinctioncoefficients and wavelengths of light

FIG. 35 is a graph illustrating a technique for identifying respiratorydisorders using the extra-thoracic monitoring system of the presentinvention.

DETAILED DESCRIPTION OF THE PRESENTLY PREFERRED EMBODIMENTS OF THEINVENTION

Mammals displace blood from within the thorax to the extra-thoraciccirculation with each heartbeat and with each breath. The displacementof blood from the thorax by the heart is due to the volumetric dischargeof blood from the heart into the arterial system, and, in particular,into the aorta. For example, the work of the heart changes the pressureof the blood in the thoracic vessels relative to that in the systemiccirculation. In addition, respiration or breathing also causes apressure change in the thorax, and, hence, displacement of blood fromthe thorax to the extra-thoracic circulation. For purposes of thepresent invention, the changes in volumetric discharge of blood from theheart that cause a measurable distention of the extra-thoracic arterialcirculation and the pressure changes in the thorax that occur duringeach heartbeat and during each respiratory cycle that also cause ameasurable distention of the extra-thoracic arterial circulation arecollectively referred to as an “intra-thoracic pressure” changes.

Volumetric flow is the result of a pressure gradient. Thus, changes inintra-thoracic pressure caused by the heart beating and the lungsbreathing are reflected by changes in pressure in the extra-thoracicarterial circulation, which is more commonly known as “blood pressure”,and also by changes in the extra-thoracic arterial circulation volume.That is, the amount of blood displaced systemically, as well as theblood pressure, changes with the amount of pressure generated within thethorax due to breathing and with the work of the heart. When there is anincrease in the intra-thoracic pressure, the displaced volume of bloodfrom the thorax and the blood pressure increases, causing theextra-thoracic arteries to increase in diameter. This phenomenon isreferred to herein as “vessel distention.”

The present invention contemplates monitoring intra-thoracic pressurechanges due to respiration or cardiac function by monitoring changes inthe extra-thoracic arterial circulation resulting from respiratory orcardiac induced intra-thoracic pressure changes. According to oneembodiment of the present invention, the patient's intra-thoracicpressure changes that are primarily due to respiration are monitored bymonitoring a characteristic of the extra-thoracic arterial circulationthat is influenced by the respiratory induced intra-thoracic pressurechanges. In essence, this embodiment of the extra-thoracic monitoringsystem provides an indirect pleural pressure monitor that effectivelyacts as a surrogate to placing an esophageal pressure monitor in thepatient. One potential application for this embodiment of the presentinvention is to monitor a patient's respiratory effort, also know aswork of breathing. The greater the respiratory effort, the greater thechange intra-thoracic pressure, which the present invention monitorsfrom the patient's extra-thoracic arterial circulation.

In another embodiment, the patient's intra-thoracic pressure changesthat are due to cardiac activity are monitored by monitoring vesseldistention in the extra-thoracic arterial circulation. In essence, thisembodiment of the extra-thoracic monitoring system of the presentinvention provides an indication of changes in blood pressure thateffectively acts as a surrogate to placing an arterial line in apatient, which is a relatively invasive procedure. Because the presentinvention allows the monitoring of the cardiac pressure changes to takeplace non-invasively and substantially continuously, specific cardiacevents, such as pulsus paradoxis, can be readily identified.

FIG. 1 is a schematic diagram of the basic components of anextra-thoracic monitoring system 30 according to the principles of thepresent invention that is capable of monitoring intra-thoracic pressurechanges due to respiration or cardiac function by examining changes inthe extra-thoracic arterial circulation associated with theintra-thoracic pressure changes. In its most basic form, extra-thoracicmonitoring system 30 includes a sensor 32, a processor 34, and aninput/output interface 36. It should be noted the monitoring system ofthe present invention is also referred to in this application as an“extra-thoracic monitoring system.”

Sensor 32 is any sensor suitable for detecting a physiologicalcharacteristic of a patient associated with pressure changes in theextra-thoracic arterial circulation and for outputting a signalindicative of such physiological characteristic. As noted above, in oneembodiment of the present invention, sensor 32 is an optical sensor thatmonitors vessel distension. It will be better understood upon reviewingthe various embodiments of the present invention discussed below, thatthe types of sensors suitable for use as sensor 32 depends on theembodiment of the invention being practiced.

Processor 34 is a processing element, such as a microprocessor, thatreceives the output from sensor 32 and processes this data to producethe desired output. For example, the respiratory monitoring embodimentof the present invention contemplates that processor 34 produces apulmonary pressure signal as a measure of the patient's intra-thoracicpressure due to respiration by isolating breath related pressurevariations from the signal from sensor 32. On the other hand, thecardiac monitoring embodiment of the present invention contemplates thatprocessor 34 produces a cardiac pressure signal as a measure of thepatient's intra-thoracic pressure due to cardiac function by isolatingcardiac related pressure variations from the signal from sensor 32. Itis to understood that processor 34 includes the necessary memory andprocessing capability to implement the features of the presentinvention.

Input/output interface 36 is any device that provides the output of theprocessor, such as the thoracic pressure signal or the cardiac pressuresignal, in a human perceivable format. In short, I/O interface 36communicates information or data between a user and processor 34.Examples of common input/output interfaces suitable for this purposeinclude a keypad, strip chart, and display. Other communicationtechniques, either hard-wired or wireless, are also contemplated by thepresent invention. For example, the present invention contemplatesproviding a smart card terminal that enables data to be downloaded fromprocessor 34 onto the smart card. Other exemplary, I/O interfaces andtechniques adapted for use with the pressure support system include, butare not limited to, an RS-232 port, CD reader/writer, DVD reader/writer,RF link, and/or modem (telephone, cable or other). In short, anyconventional technique for providing, receiving, or exchanging data withthe processor are contemplated by the present invention as input/outputinterface 36.

I. SYSTEM ARCHITECTURE

As noted above, one embodiment of the present invention contemplatesmonitoring intra-thoracic pressure changes due to respiration bymonitoring changes in the extra-thoracic arterial circulation resultingfrom the respiratory induced intra-thoracic pressure changes. Thus, forthis embodiment, processor 34 receives the output of sensor 32 andproduces a pulmonary pressure signal as a measure of a patient'sintra-thoracic pressure due to respiration by isolating breath relatedpressure variations in the first signal. This pulmonary pressure signalis provided to input/output device 34.

FIG. 2 illustrates one embodiment of extra-thoracic monitoring system 30in greater detail, including a housing 40, which contains processor 34and input/output interface 36, a respiratory sensing system 35, and anoptical system 42. The respiratory and optical systems are suitable foruse as sensor 32 from FIG. 1. FIG. 3 schematically illustrates a firstembodiment for the components of pulmonary monitoring system 38.Referring to FIG. 2, housing 40 includes a display area 44, such as anLED, LCD, or any other conventional display, a speaker 46, and aconnection terminal 48, as examples of output components of theinput/output interface. Housing 40 includes a keypad 50 as an example ofan input component of the input/output interface. Of course, as notedabove, any conventional input/output interface, such as a touch screenor wireless communication device, is contemplated by the presentinvention as being suitable for interfacing a user or other peripheraldevice with the processing elements of the monitoring system.

Optical system 42 includes a photoemitter 52, or a plurality ofphotoemitters, adapted to transmit light through a portion of a patient54, and a photodetector 56 adapted to receive light after having beenpassed through the patient, as indicated by arrow A. The processingelements in housing 40 that control emitter 52 and receive the signalfrom photodetector 56 communicate with the emitter and detector viacommunication lines 58. The present invention contemplates anyconventional technique for communicating between the processor andphotoemitter 52 and between the processor and photodetector 56. AlthoughFIG. 2 shows the light being passed through a patient's finger, thepresent invention contemplates passing the light through anyextra-thoracic portion of the patient where the arterial circulation canbe monitored, such as the ear, toe, or nasal septum. Optical system 42in combination with the processing elements constitutes aphotoplethysmographic monitoring system that, in this embodiment,monitors the changes in vessel distention resulting from the respiratoryaction of the patient, spontaneous or otherwise.

In the illustrated embodiment, the optical system is a transmissive typeof photoplethysmographic monitor. It is to be understood that thepresent invention also contemplates that the optical system is areflective type of photoplethysmographic monitor in which light isdirected into the tissue and the photodetector detects the lightreflected back out of the patient.

In an exemplary embodiment of the present invention, vessel distentionis determined by passing light through the finger or other appendage. Asnoted above, a change in blood volume causes a change in vesseldiameter. The absorbance of light increases and decreases as the vesseldiameter increases and decreases. The present invention measures thechange in vessel distention by continuously measuring the change inabsorbed light.

It should be noted that the present invention contemplates using areflectance type of photoplethysmographic in place of the transmittancetype sensor shown in the figures and described herein. In a reflectancetype photoplethysmographic sensor, light is directed into the patientand the amount of light reflected back from the patient is monitored andbecomes the photoplethysmographic signal. A transmittance type sensor isbelieved to provide a stronger signal than a reflectance type sensor.

As noted above, photoemitter 54 in optical system 42 delivers lightthrough a portion of the patient and measures the light passingtherethrough. The light signal received by detector 56 in optical system42 is provided to the processor where it is processed in real time toseparate breathing from heartbeats.

FIG. 3 schematically illustrates the components of a first embodiment ofan optical sensor based pulmonary monitor 38, including the componentsfor processing the raw photoplethysmographic signals 60 from detector56. An example of photoplethysmographic signals 60 is shown in FIG. 10.Photoemitter 52 preferably emits light at a wavelength that isrelatively insensitive to an oxygen saturation level of the patient,while relatively sensitive to changes in path length between thephotoemitter and photoreceiver. That is, it is preferred that changes inoxygen saturation do not adversely influence the vessel distensionmeasurement. This provides an advantage in that the oxygen saturation ofthe patient (SpO₂) does not have to be factored in when determining thechange in vessel distention Δd. Thus, a single LED, operating, forexample, and approximately 800 nm, is capable of performing the vesseldistension monitoring.

Respiratory monitoring sensor 35 provides a signal indicative ofrespiratory rate (f_(RR)) or breathing frequency (f_(breathing)). Thisinput is needed in some of the embodiments of the present invention tocalculate the vessel distension. The present invention contemplates thatany device that is capable of identifying the respiratory cycles of thepatient can be used as respiratory monitoring sensor 35. For example, aneffort belt provides a relatively good quality respiratory signal. Otherembodiments of the present invention contemplate using a device thatinterfaces with a patient's airway to measure the pressure or flow attheir airway, such as a nasal canula, mask, or flow sensor (pneumotach).An example of a suitable respiratory monitor is disclosed, for example,in U.S. Pat. No. 6,544,192.

FIG. 3 is a more detailed schematic diagram of a first embodiment of anextra-thoracic monitoring system 30 according to the present invention.Monitoring system 30 operates using a microcomputer system withsatellite sub-systems. A main computer system 62 utilizes conventionalprocessing techniques and applications, such as Windows 2000®, VisualBasic, and Softwire, to process the received data, display results, andcommunicate to a interface module 64. Monitoring system 30 includes apressure card 66 and a photoplethysmograph (SpO₂) module 68. Pressurecard 66, photoplethysmograph module 68, interface module 64, andprocessor 62 define the major components of the processing system. In apreferred embodiment, the pressure card 66 and a plethysmograph (SpO₂)module 68 are provided as separate circuit boards to provide separateconnections to the associated patient interface devices.

Interface module 64 receives and transmits analog and binary signalsused to control the SpO₂ module and to convert those signals to digitalvalues. In an exemplary embodiment of the present invention, pressurecard 66 has four pressure sensors that physically connect to arespiratory monitoring device 35, such as a patient interface device 67,e.g., mask, nasal canula, pneumotach (differential pressure sensor thatmeasures a pressure difference across a flow restriction). The pressuresensors on pressure card 66 that are connected to the patient interfacedevice will depend on the patient interface device being used. Theoutputs of the sensors are scaled to the input voltage of the interfacecard for maximum voltage signals at maximum required pressure range.Computer 62 uses the digital values of this signal and calibrates it todisplay proper levels. The pressure card receives pressures to monitorthe following parameters:

O₂ flow (oxygen supplied to the patient);

Patient airway pressure (the pressure of the patient airway);

Patient flow (monitored using two pressures from a pneumotach); and

Patient flow measured from a cannula or mask.

The SpO₂ Module is used normally to read the SpO₂ value or non-invasiveblood O₂ gas of the patient. The device also conveys delta signalsindicative of respiratory effort and cardiac function as well as the %Pulse Paradox at the finger probe.

The SpO₂ Module connects to a patient interface device, such as opticalsystem 42.

The operation of the electronic components of cardio/pulmonarymonitoring system 30 according to an exemplary embodiment of the presentinvention will now be described with reference to FIGS. 4-7. FIG. 4 is aschematic diagram of the hardware components of the SpO₂ module 68 ofthe cardio/pulmonary monitoring system of FIG. 3, and FIG. 5 is adetailed circuit diagram corresponding to the schematic diagram of FIG.4. FIGS. 6A-6K illustrate portions of the circuit shown in FIG. 5, andFIG. 7 is a timing diagram for the operation of the circuit shown inFIG. 5. It is to be understood that FIG. 4 and the description of theoperation of the embodiment of the invention shown in FIG. 5 andpresented below is intended to describe only the operation of theelectronics components of the system. Details as to how the systemprocesses the output of the SpO₂ module are discussed in other portionsof this application.

The description of the operation of the circuit begins at finger probe70, which also corresponds to optical system 42. Probe 70 preferably hastwo LED's, a Red 72 and IR (Infrared) 74. Light from the two LED's istransmitted into the finger at timed intervals. A sensor 56 in the formof a photo cell in the probe monitors the light transmitted through thefinger. A timing control circuit 78 controls the operation of the LEDs.For example, there is a time period in which both LED's are turned offto measure the ambient light around the sensor. This ambient light cancause problems with the signals and needs to be removed.

Sensor 56 transmits the different levels of the signal created by eachlight source. (LED's/Ambient light). Transimpedance differentialamplifiers 80 amplify this signal and then output it to an ambient lightcanceling circuit 82. See FIG. 6A.

FIG. 6B illustrates an ambient light canceling circuit 82 that receivesthe output of transimpedance differential amplifiers 80. The ambientlight canceling circuit operates as follows: when timing control circuit78 has both the Red and IR LED's off, the ambient light is the onlylight the sensor has for an output. The Ambient light is sampled, andthe value of the signal is held in a capacitor 84 tied to ground using aFET 86. When the FET is turned off, the value stored in the capacitor isused in the path of the Red and IR signal string. This stored value inthe capacitor removes the error of the ambient light. The out put signal88 now has a true value of the needed signals.

Now that the ambient light is removed from the circuit, the sensorsignal needs to be adjusted for the user. This is done in a primary gainstage 90. See FIG. 6C. In a presently preferred configuration, thecircuit can be adjusted between three different gains Low, Medium, andHi. Primary gain stage 90 allows the circuit to accommodate patientswith small hands or patients whose extremities are cold. In thesesituations, for example, too much light will be transmitted to thesensor, causing the LED in the finger probe to be forced. To preventthis, a low gain is used for the patients with small or cold hands, sothat the LEDs can operate in the proper range. Detecting that the gainneeds to be adjusted and providing the right gain can be accomplishedautomatically, via the main computer. The present invention alsocontemplates controlling these gains manually.

The signal from the primary gain stage is ready to be split between theRED and IR circuits. This splitting is accomplished in the presentinvention using a sample/hold circuit 96 a and 96 b. See FIGS. 6D and6E. When the timing for the RED LED is on, a sample is taken of thesignal from the sensor, and it is stored on a capacitor during the REDtiming cycle. The same happens during an IR cycle, holding the value intime of the IR level from the sensor. Like points in a chart, the sampleand hold circuits create the pulse plethysmograph waveforms that will beused to measure SpO₂. The signal from the RED LED is less than the IRsignal. As a result, a gain stage is added to the Red S/H circuit.

The DC level of the signals are filtered using Low Pass Filters (LPF)100 a and 100 b, e.g., a 19.9 Hz LPF, to remove switching noise from thesample/hold circuits. See FIG. 6F. The outputs 101 a and 101 b of thelow pass filters are connected to the analog input CH03-pin8 (red) andCH04-pin10 (IR) of the main computer, as well to a differentialamplifier U3 102 a and 102 b. See FIG. 6G.

Each of the two Differential amplifiers (Red and IR) 102 a and 102 bused in this stage has two input signals, the output of the sample/holdfilter 101 a and 101 b, and the Reference set by the main computer'sanalog output A0 (IR) 104 a and A1 (RED) 104 b. There are two functionsfor the analog outputs of the main computer. The first is to set up thelight intensity. The main computer uses the differential amplifier tocontrol the brightness of the finger probes LEDs. When the light iscorrect, the DC valve at the input of the differential amplifier equalsthe commanded analog signal of the main computer. The output of thedifferential amplifier will equal zero.

The other operation of the analog output of the computer is to keep theoutput of the differential amplifiers 102 a and 102 b inside a window ofoperation without adjusting the finger probes light intensity. Thisoperation will lock out the light controller, but the voltage of theanalog output will change to keep the output of the differentialamplifier to zero if it should fall outside of a window range. This islike an AC coupled circuit without the delays of a capacitor and theability to know the value of the change made.

As shown in FIGS. 4 and 5, outputs 106 a and 106 b of differentialamplifiers 102 a and 102 b accomplish two functions: a light controlleroperation, and an AC gain and filter operation, referred to as the ACsignal conditioning, to provide the input to the main computer.

During the light adjustment for brightness of the LEDs (the lightcontroller operation), the error voltage of the DC signals (DCX (Red orIR) and the DC Ref (from computer for the red or IR)) is compared in arespective comparing circuit 110 a or 110 b. See FIG. 6H. The differenceout of the differencing amplifier is the error voltage needed to correctthe LED intensity. During this needed change of brightness, the maincomputer closes the switch to the integrator, causing it to adjust thecontrol voltage of the drive circuit so that the DC value of the red orIR LED's will equal the reference voltage. At the end of this operation,the computer opens the switch and controller will freeze with the lastvalue out.

The output of the comparing circuit 110 a and 110 b is provided to aProportional level shift gain/mixing circuit 112. See FIG. 6I. Tounderstand the operation of this circuit, it should first be noted thatthe LED drive circuits 78 need a voltage level of 0-1 volt to operatethe LEDs. The drive circuit also needs the voltage of both LEDs set intoa timing string to match its output between the RED and IR circuits. Themixing part of the circuit uses two analog switches to select betweenthe two integrators, making a control string for the drive circuit.Because the output of the integrator is between −5 to +5 volts, the gainwill need to be divided by 10 to provide the needed plus and minus 0.5volts. The circuit will then level shift that voltage by 0.5 voltsgiving the needed 0-1 volt for the next stage.

The final stage 114 of the control loop (the light controller operation)is in two parts: the voltage-to-current converter, and the invertingoutput stage. See FIG. 6J. The current converter uses a 10-ohm sampleresistor in the path of the LEDs output drive circuit. The voltage dropacross the resistor should equal 10× the current or 1 Volt will equal100 mA. Using this voltage and comparing it to the output of the controlvoltage, the amplifier will adjust its output voltage to adjust thecurrent to be equal to the commanded current string.

Part two of the drive breaks down the string so that it can control theproper LED to the timing used in the other circuits. In the illustratedexemplary embodiment, NPN transistors (2N3904) (X4-X7) are used to setthe current. FETs 116 a and 116 b (2N7002) at the base of thetransistor, when turned on, will shut down the current of that leg ofthe H Bridge drive circuit. The PNP transistors at the top of the HBridge drive (2N3906) also are used to shut down or enable the currentto flow in the proper direction with the proper timing. For example,during the Red Drive time, the control voltage is 0.6 volts (60 mAcommand RED timing) RB6 will go low RB7 will remain high. Because RB6 islow, the left 2N3906 is on, allowing current to pass in that leg. Theright transistor is off if RB7 is high, and no current will flow in theother leg. With RB6 low, the right 2N7002 FET will be turned off, andthe control voltage of the U6 amplifier will feed the right 2N3904transistor. The current then flows from 5VDC down the left leg PNPtransistor into the probes RED LED down the right NPN transistor andinto the 10-ohm current sample resistor for the feedback for currentcorrection of the U6 amplifier. The operation is the same for the IRLED, only RB6 is high and RB7 is low, and the other path of the H driveis used.

As noted above, the outputs of differential amplifiers 102 a and 102 bare provided to AC gain circuits 120 a and 120 b and filter circuit 124a and 124 b to perform AC signal conditioning on the Red and IR signals.This is done prior to providing these signals to the main computer.These operations are discussed below with reference to FIG. 6K.

Signals from differential amplifiers 102 a and 102 b, which originatedfrom the probe, contain information that can determine SpO₂, heart rate,and other information about the person wearing the probe. The signalsalso have high frequency noise, low frequency offsets, and low or highsignal levels. The conditioning circuits can be broken down into threeparts: (1) differential offset control, (2) gain circuits, and (3) lowpass filter. Each of these three parts are discussed in turn below.Conditioned waveforms for the RED and IR are then sent to the maincomputer via interface module 64.

Because the system of the present invention does not use a high passfilter in the hardware, offsets can cause signals to go into theoperating rail. The computer looks for this offset and adjusts the Ref.Voltage. (Note: this voltage is entered into the previous stage). Thislets the computer adjust the offset correctly. Knowing the value of theoffset, the data internal to the computer, can be spliced. Because thecomputer works in a virtual world it can simulate voltages outside ofthe range in which the circuit can operate.

Gain stages 120 a and 120 b are used to adjust the gain or peak-peakvalue of the signal to keep it inside a window of operation. Too muchgain and the waveform will run out of voltage, distorting the peeks ofsignal. Too little, and the final signal will not be able to resolve thevalues of the waveform, reducing the accuracy of the system. In thisstage, the main computer will provide the correct window of operation byadjusting the gain.

Low pass filters 124 a and 124 b are used to pass signals less than thechosen cutoff frequency, blocking noise and other interference signals.The 19.9 Hz low pass used has a cutoff frequency of just above the heartrate range. This means noise variations above 19.9 Hz, or frequenciesabove the breathing range, are reduced/removed from the waveforms thatare being conditioned.

Timing for the circuit is important. A small PIC controller is used toprovide the timing signals RB1 to RB7 (output pin names of the PICcontrol). FIG. 7 is a raw diagram of the timing and waveform creationmethod.

Starting with RB5 in the timing chart shown in FIG. 7 shows that theambient light sample starts the timing for each cycle of the PICcontroller. The dots illustrate when the error voltage created by theambient light is stored. Also note that the Red and IR LEDs are off. Thesignal string from the probe also will show a lower or no signal, asindicate at point A at the lower end of the timing chart. (Signalwaveform string is for theory only.)

The next phase of the timing cycle, is the RED timing. Setting RB6 lowturns on the transistor circuit used in the drive stage to allow currentthrough the RED LED. Next, the correct current command is passed to theoutput circuit by setting RB1 high. The LED turns on and lightstabilizes almost instantly. A delay is set to insure the LED is on andstable then RB2 is turned on sampling the waveform (see point C of thelower portion of the timing chart showing part of the waveform that isbeing sampled). When RB2 goes low, the voltage value is stored on thehold capacitor for the red (point E of lower cart shows one of many heldvoltage points). RB1 and RB6 go to their output off for the Red LED.

During the IR phase of the timing cycle, the transistor circuit used inthe drive current through the IR LED is turned on by setting RB7 low.Next, the correct current command is passed to the output circuit bysetting RB3 high. The LED turns on and light stabilizes almostinstantly. A delay is set to insure the LED is on and stable then RB4 isturned on sampling the waveform (point D of the lower chart of FIG. 7showing part of the waveform that is being sampled). When RB4 goes low,the voltage value is stored on the hold capacitor for the red (point Fof lower cart shows one of many held voltage points). RB3 and RB7 go totheir output off status for the Red LED.

II. NON-INVASIVE VESSEL DISTENSION MONITORING

As noted above, the present invention contemplates monitoring thepatient by non-invasively monitoring the patient's vessel distension(NIVD). This is accomplished by measuring the intensity of lighttransmitted through the patient's tissue. FIG. 8 is a schematic diagramof the user's tissue disposed between a photoemitter 52 and aphotodetector 56 of the present invention. This figures illustrates thetissues that affect the absorbance of light passing through the tissues.It can be appreciated that there are two components associated with theabsorbance of light: a DC component that does not change with eachheartbeat or breath; and an AC component Δd that is the result ofarterial blood pulsates due to the heart beating or breathing. FIG. 8illustrates only changes in path length due to the cardiac activity. Itcan be appreciated that changes in path length due to respiration willbe similar, except that d_(min) (I_(H)) will coincide with the lowestpressure occurring through one respiratory cycle, and d_(max) (I_(L))will coincide with the highest pressure occurring through onerespiratory cycle.

The present inventors determined that the change in path length Δd canbe determined if S_(P)O₂ and the concentration of total functionalhemoglobin are known. The change in path length, Δd, due to breathing,which is referred to as the Thoracic Δd or NIVD_(Thoracic), and thechange in path length to the heart stroke volume, which is referred toas the Cardiac Δd or NIVD_(Cardiac), are determined using the followingformula:

$\begin{matrix}{{NIVD} = {{\Delta \; d} = {\frac{- {\ln \left( \frac{I_{X}}{I_{H}} \right)}}{\begin{Bmatrix}{\left\lbrack {\left( ɛ_{{HbO}_{2}} \right){\left( \lambda_{IR} \right) \cdot \frac{{SpO}_{2}}{100}}} \right\rbrack +} \\\left\lbrack {{ɛ_{Hb}\left( \lambda_{IR} \right)} \cdot \left( {1 - \frac{{SpO}_{2}}{100}} \right)} \right\rbrack\end{Bmatrix}c_{TotHb}} \cdot}}} & (1)\end{matrix}$

Where:

I_(X) is the intensity of light transmitted through the tissue at anygiven time;

I_(H) is the peak intensity of light transmitted through the tissue atany given time;

ε_(HbO2) is the extinction coefficient for oxygenated hemoglobin, i.e.,functional hemoglobin that is fully saturated with oxygen;

ε_(Hb) is the extinction coefficient for reduced hemoglobin, i.e.,functional hemoglobin that is not fully saturated with oxygen;

λ is the wavelength of light being directed into the user; and

c_(TotHb) is the total concentration of functional hemoglobin(c_(Hb)+C_(HbO2)).

The total concentration of functional hemoglobin in whole blood(c_(TotHb)) is given by a number of sources, some to varying degrees ofaccuracy. The extra-thoracic monitoring system of the present inventionassumes a total concentration of 2.265 millimole per liter (mM/L), basedon a patient with normal amounts of dyshemoglobin. Other errors exist.For example, cigarette smoking temporarily “steals” small amounts ofhemoglobin, creating dyshemoglobins that absorb light differently. Bylimiting the degree of accuracy of the total concentration of hemoglobinto 2.265 mM/L, any actual changes are so small they are most likelyinsignificant.

The other parameters in Equation (1) are either computed (SpO₂), orgiven (λ). In a presently preferred embodiment, an average SpO₂, such asan average SpO₂ over 2.5 seconds, is used in equation (1) to prevent asingle errant event from providing unreliable results. The method forfinding I_(L)/I_(H), and over what time frame, becomes the maindifference between the two different vascular distension measurements.All NIVD measurements are converted to micrometers by multiplying theresult of equation 11 by 10,000.

Cardiac Δd (NIVD_(Cardiac)) is a measure of the change in path lengthfrom one heartbeat to the next, and, by normalizing this signal over abreath, generates a percent change. Equation (1) is used to determineCardiac Δd by letting I_(X) correspond to the intensity of lighttransmitted through the tissues at any given time, x, during one cardiaccycle, and by letting I_(H) correspond to the peak intensity of lighttransmitted through the tissues during one cardiac cycle. As a result,Cardiac Δd represents the change in diameter of the arterial vesselsfrom their minimum value (diastole) to their value at time, x, duringone cardiac cycle.

Thoracic Δd (NIVD_(Thoracic)) is a measure of the effect of thoracicpressure swings on the effective path length seen at the probe site.Thoracic Δd is an alternative to the awkward, invasive conventionaltechnique of swallowing an esophageal balloon catheter in order tomonitor thoracic pressure swings. Equation (1) is used to determineThoracic Δd by letting I_(X) correspond to the intensity of lightpassing through the tissues at any given time, x, during one breath, andby letting I_(H) correspond to the peak intensity of light passedthrough the tissues during one breath. The diameter of the arterialvessels are at a minimum when the lung pressure is at atmosphericpressure and ignoring any effect on vessel distention due to cardiacfunction. As a result, Thoracic Δd represents the change in diameters ofthe arterial vessels from their minimum value to their value at time, x,during one breath;

III. OPERATION OF THE SYSTEM

FIG. 9 is a schematic diagram of an exemplary embodiment of theextra-thoracic monitoring system 30 according to the principles of thepresent invention. Signals 60 from photodetector 56 are provided to anA/D converter, which is equivalent to interface module 64 of FIG. 3.FIG. 10 illustrates a raw hypothetical plethysmograph signal 60 detectedby the photodetector, which forms the plethysmograph monitoring portionof the extra-thoracic monitoring system and is equivalent to SpO₂ module68 and patient interface 42 of FIG. 3. A respiratory sensor 35 monitorsa characteristic of the patient indicative of respiration, such asairflow, and provides a respiratory signal 63 indicative thereof. FIG.11 is a graph of respiratory signal 63, which is also a component of theraw plethysmograph signal of FIG. 10.

The signal from the SpO₂ module and the pressure card (airflow sensor35) are provided to buffer 150 for use by the main computer and insubsequent processing steps. The plethysmograph signal 60 is provided toa Fast Fourier Transform operator 152. An example of the resultingoutput frequency spectrum signal 154 from Fast Fourier Transformoperator 152 is shown in FIG. 12A. If necessary, an offset is removed bya zero out DC component module 156 to produce a frequency spectrumsignal 158 without any DC bias. See FIG. 12B.

The peaks of the frequency spectrum signal 158 are detected by peakdetector 160. Detecting the peaks is necessary to select the properfiltering frequencies to be applied to plethysmograph signal 60 indynamic filter 162.

The present invention contemplates that the Cardiac Δd or the ThoracicΔd can be monitored using the extra-thoracic monitoring system of thepresent invention. The determination of which one of these variables (orboth) is to be monitored is based on the filtering applied to theplethysmograph signal by dynamic filter 162.

Determining the vessel distention due to respiration (NIVD_(Thoracic))involves isolating the respiratory rate frequency component (f_(RR))from the frequency components of frequency spectrum signal 158, which isaccomplish by dynamic filter 162. The plethysmograph signal 60 is thenfiltered so as to isolate the respiratory rate frequency componentf_(RR), thereby producing a vessel distention signal which is asurrogate for an intra-thoracic pressure measurement.

In one embodiment of the present invention, the peak in the frequencyspectrum signal 158 corresponding to the breathing rate is detectedbased on the monitored respiratory rate from sensor 35, which isdetermined, at least in part by peak detector 163. Once the respiratoryrate frequency component f_(RR) is identified, a cutoff frequencyf_(cutoff) is determined. See FIG. 13A. In this embodiment, if therespiratory rate frequency component f_(RR) is less than the heart ratefrequency component f_(HR), the dynamic filter sets the cutoff frequency(f_(cutoff)) as f_(RR)+f_(smear). Filter 162 then low pass filtersplethysmograph signal 60 at the cutoff frequency f_(cutoff). If therespiratory rate frequency component f_(RR) is deemed to be greater thanthe heart rate frequency component f_(HR), the dynamic filter sets thecutoff frequency (f_(cutoff)) as f_(RR)−f_(smear). See FIG. 13B. Filter162 then high pass filters plethysmograph signal 60 at the cutofffrequency f_(cutoff). In this embodiment, f_(smear) is a predeterminedthreshold frequency.

In another embodiment of the present invention, the peak in thefrequency spectrum signal 158 corresponding to the heart rate isdetected using conventional techniques, such as that used in pulseoximetry or in EKG monitoring. This is determined, at least in part, bypeak detector 163. Of course, other techniques for detecting heart ratecan be used. Once the heart rate frequency component f_(HR) isidentified, a cutoff frequency f_(cutoff) is determined in a mannersimilar to that discussed above. More specifically, if the respiratoryrate frequency component f_(RR) is less than the heart rate frequencycomponent f_(HR), the dynamic filter sets a cutoff frequency(f_(cutoff)) as f_(HR)−f_(smear). Filter 162 then low pass filtersplethysmograph signal 60 at the cutoff frequency f_(cutoff). If therespiratory rate frequency component f_(RR) is greater than the heartrate frequency component f_(HR), and sets the cutoff frequency(f_(cutoff)) f_(RR)+f_(smear). The filter then high pass filtersplethysmograph signal 60 at the cutoff frequency f_(cutoff). In thisembodiment, f_(smear) is a predetermined threshold frequency.

The present invention also contemplates that the size of f_(smear) canbe adjusted. An example of this is discussed below. In any event,whether or not f_(smear) is adjustable or fixed, f_(smear) should belarge enough to ensure that sidelobes contributing to the main peak inthe frequency band of interest are captured, so that adequate filteringis performed.

Determining the vessel distention due to cardiac function(NIVD_(cardiac)) involves isolating the heart rate frequency component(f_(HR)) from the frequency components of frequency spectrum signal 158,using a process similar to that discussed above. As with the previousembodiment, the cutoff frequency can be determined from the respiratoryrate frequency component f_(RR) or from the heart rate frequencycomponent f_(HR).

If the respiratory rate frequency component f_(RR) is known, and if therespiratory rate frequency component f_(RR) is less than the heart ratefrequency component f_(HR), the dynamic filter sets the cutoff frequency(f_(cutoff)) as f_(RR)+f_(smear), thereby isolating the heart ratefrequency component f_(HR). Filter 162 then high pass filtersplethysmograph signal 60 at the cutoff frequency f_(cutoff). If therespiratory rate frequency component f_(RR) is deemed to be greater thanthe heart rate frequency component f_(HR), dynamic filter 162 sets thecutoff frequency (f_(cutoff)) as f_(RR)−f_(smear). See FIG. 13B. Filter162 then low pass filters plethysmograph signal 60 at the cutofffrequency f_(cutoff). In this embodiment, f_(smear) is a predeterminedthreshold frequency.

If the heart rate frequency component f_(HR) is known, and if therespiratory rate frequency component f_(RR) is less than the heart ratefrequency component f_(HR), the dynamic filter sets the cutoff frequency(f_(cutoff)) as f_(HR)−f_(smear), thereby isolating the heart ratefrequency component f_(HR). Filter 162 then high pass filtersplethysmograph signal 60 at the cutoff frequency f_(cutoff). If therespiratory rate frequency component f_(RR) is deemed to be greater thanthe heart rate frequency component f_(HR), dynamic filter 162 sets thecutoff frequency (f_(cutoff)) as f_(HR)+f_(smear). See FIG. 13B. Filter162 then low pass filters plethysmograph signal 60 at the cutofffrequency f_(cutoff). In this embodiment, f_(smear) is a predeterminedthreshold frequency.

The present invention also contemplates using a band-pass filtercentered on the frequency component of interest to isolate thatcomponent from the frequency spectrum signal. For example, if therespiratory rate frequency component f_(RR) is known and theNIVD_(Thoracic) is being monitored, an upper cutoff frequency for theband pass filter can be set as f_(RR)+f_(smear) and a lower cutofffrequency for the band pass filter can be set as f_(RR)−f_(smear).Filter 162 then band pass filters plethysmograph signal 60 at theseupper and lower cutoff frequencies. If the heart rate frequencycomponent f_(HR) is known and the NIVD_(Cardiac) is being monitored, anupper cutoff frequency for the band pass filter can be set asf_(HR)+f_(smear) and a lower cutoff frequency for the band pass filtercan be set as f_(HR)−f_(smear). Filter 162 then band pass filtersplethysmograph signal 60 at these upper and lower cutoff frequencies.

The present invention contemplates that the cutoff frequency isrecalculated for each breath, thereby providing a very fast response toany changes in the patient, which provides a more accurate measurementof NIVD_(Thoracic) or NIVD_(Cardiac). Of course, the cutoffs can also becalculated less frequently or the frequency by which the cutoffs arerecalculated can be determined based on the monitored condition of thepatient, thereby maximizing system efficiency. For example, if thepatient is relatively stable, the cutoffs can be recalculated lessfrequently than when the patient is not.

An example of the NIVD_(Thoracic) output waveform 191 and NIVD_(Cardiac)output waveform 193 produced by the present invention is shown in FIG.14. This figure also shows a total or raw NIVD signal 60. The output ofdynamic filter 162 (NIVD_(Thoracic) signal 191 and/or NIVD_(Cardiac)signal 193), which is the filtered plethysmograph signal, can be used ina variety of ways. For example, the present invention contemplatesremoving any phase shifts, if desired, via a phase shift removalcomponent 180. The signals can be displayed, transmitted, stored, oroperated upon in any manner by controller 188 and input/output device190. It should be noted that the present invention contemplates thateach component 150-180 operates under the control of controller 188.Arrows 192 signify this aspect of the invention.

It can be appreciated that other physiological characteristics thatmanifest themselves as a pressure or volume change in the patient'sarterial circulatory system can be monitored by the extra-thoracicmonitoring system of the present invention. This is accomplished bysetting the frequency cutoff to select or isolate the frequenciesassociated with these characteristics in the frequency spectrum signal.For example, Burton's Waves or Traub-Herring Waves, which are relativelyslow changes in the patient's circulatory pressure or volume, can bedetected by selecting the frequency cutoff to remove the higherfrequency signals, such as breathing and heart rate.

The two primary parameters for calculating blood flow from the heart arestroke volume and heart rate. Because blood flow equals stroke volumetimes the heart rate, the flow changes as a result of altering strokevolume, heart rate, or both. In order to assess the change in bloodflow, the present invention contemplates plotting the heart rate andpathlength changes that occur for each heartbeat within each breath.

IV. PHYSIOLOGICAL ROLL-OFF

The inventors became aware that the raw NIVD signal, as well as theNIVD_(Thoracic) or NIVD_(Cardiac) signal, are attenuated as thepatient's breathing frequency increases. The “roll-off” of the raw NIVDsignal's peak-to-peak values as the breathing frequency increases isshown by line 159 in FIG. 15. Line 159 is a trend line for thepeak-to-peak NIVD measurements taken over a range of breathing rates.The present invention contemplates using this roll-off to correct orcompensate for the raw NIVD signal (Δd), NIVD_(Thoracic) (Cardiac Δd),NIVD_(Cardiac), (Thoracic Δd) signals, or any combination thereof basedon the monitored breathing rate. For example, once an NIVD value isdetermined, the patient's breathing frequency at that time is alsodetermined from respiratory sensor 35. The NIVD peak-to-peak value canthen be corrected based on the known NIVD_((peak-to-peak)) versusbreathing frequency relationship F. This correction can be done usingany conventional signal processing technique. For example, a look-uptable can be generated and used to provide a breathing frequencycorrection factor or an equation can be determined that represents therelationship between a measured NIVD value and breath rate. Such anequation would correspond, in general, to line 159.

It should also be noted that the resistance (R) and compliance (C) ofthe circulatory system (from the thorax to the location of theplethysmography sensor) can be determined empirically, for examplethrough the testing of a number of patients, or can be estimated usingstandard indices, such as pulse transit time. If R and C are known, theNIVD versus breathing frequency relationship can be determinedbeforehand and used to correct the NIVD value for the measured breathingfrequency.

V. INSPIRATORY TO EXPIRATORY RATIO

The inventors also became aware that a patient'sinspiratory-to-expiratory (I:E) ratio impacts frequency spectrum signal158. Namely, the present inventors determined that a decrease in the I:Eratio introduces additional harmonics in the frequency spectrum signalproduced by the FFT. This phenomena is shown in FIGS. 16-19. Thesefigures illustrate the additional harmonic frequencies, generallyindicated at 170, that appear in the frequency spectrum signal as theI:E ratio goes from 1:1 (FIG. 16) to 1:4 (FIG. 19). It should be notedthat certain patient populations have an I:E ratios that deviate from a1:1 ratio. In fact, the I:E ratio for a normal individual is in therange of 1:2-1:3. However, for some people, such as patients withasthma, Pickwickian syndrome, congestive heart failure, pulmonaryfibrosis, pneumonia or experiencing a drug overdose, the I:E ratio hasbeen determined to have an I:E ratio much less than 1:1.

Knowing that harmonics in the frequency spectrum signal near thebreathing frequency f_(RR) are created as the I:E ratio deviates from1:1, the present invention contemplates accounting for these additionalharmonics in setting the cutoff frequency. For example, in determiningThoracic Δd (NIVD_(Thoracic)) using f_(RR) as the base point, the valuefor f_(smear) can be increased as the I:E deviation from 1:1 decreases,assuming that f_(RR)<f_(HR). Using f_(HR) as the base point, the valuefor f_(smear) can be decreased as the I:E deviation from 1:1 decreases,again assuming that f_(RR)<f_(HR).

VI. INTRA-BREATH CHANGES IN HEART RATE

Blood is delivered from the heart to the systemic circulation in pulses.The average amount of blood flow leaving the heart within each cardiaccycle is known as cardiac output. Cardiac output is the product of thevolume of blood leaving the heart with each ejection portion of thecardiac cycle and the rate at which the heart is beating. Thisrelationship can be summarized as follows:

Total Average Blood Flow=Cardiac Output (mL/min)=Heart Rate(beats/min)×Stroke Volume (mL/beat).

Similarly, the average blood flow to an appendage, such as a finger, isthe product of the volume of blood delivered to the appendage withineach cardiac cycle (a.k.a. pulse volume) and the rate at which it isdelivered (pulse rate). This relationship can be summarized as follows:

Average Blood Flow to Appendage (mL/min)=Pulse Rate (beats/min)×PulseVolume (mL/beat).

As described above, a change in vessel distention arises due to a changein blood flow. The extra-thoracic monitoring system of the presentinvention provides the ability to illustrate changes in vesseldistention of an appendage, such as a finger. Arterial vessel distentionhappens during each heart cycle and each breath cycle. The averagechange in distention that occurs throughout each breath is the result ofthe average vessel distention that happens during each of the heartbeats that take place within the each breath. Because the average bloodflow to an appendage is a product of the average pulsed volume of blooddelivered and the pulse rate, the extra-thoracic monitoring systemprovides the ability to view and plot changes in pulse rate in additionto changes in distention.

FIG. 20 depicts the extra thoracic monitoring system's ability todisplay the mean change in distention that occurs within each breath(waveform 174), as well as the changes in distention that occur withineach cardiac cycle (waveform 176) and the change in cardiac cycle rate(waveform 178). Waveform 180 in FIG. 20 illustrates the raw NIVD (Δd)signal, which includes both the Cardiac Δd and Thoracic Δd components,and, thus, represents a total NIVD waveform. As can be seen in thisillustration, in this patient, for example, the distention that occursduring each cardiac cycle remains relatively the same throughout thethree breaths. However, the heart rate varies quite a bit. During thefirst breath, the heart rate varies from 1.61 to 1.75 Hz (approximately8.4 beats/min variation). During the second breath, the heart ratevaries by approximately 10.8 beats/min, and, during the third breath,the heart rate varies by approximately 12.6 beats/min.

VII. FURTHER PROCESSING

As shown in FIG. 14, if the mean vessel distention of the arterialvessels that occurs due to each heartbeat is plotted over the course ofat least one breath, the peak minus the valley of that mean (distance P)is a reflection of the pleural pressure change that occurred for thegiven breath(s). The present invention contemplates producing waveform191 from the raw NIVD signal 60 using the signal processing techniquesdiscussed above. An alternative embodiment of the present invention alsocontemplates producing waveform 191 by passing total or raw NIVD signal60 through a low pass filter. Any conventional peak-to-peak detectiontechnique can be used to determine distance P.

Another method of finding the mean vessel distention is to read a largenumber of samples of total signal 60 into an array. Then, the FFT isused to determine the heart rate. Once the heart rate is know, thepresent invention contemplates dividing the number of samples in thearray by the heart rate to create sub-arrays. Each resulting sub-arrayholds one period's worth of heart beats. Next, simply determine the meanfor each sub-array and then plot the mean of each sub array. The plotwill produce a waveform similar to that of signal 191. The advantagethat such a digital filter offers over finding the mean for eachheartbeat is improved resolution. Finding the mean for each heartbeatwill only produce one point per heartbeat, whereas the digital filterwill produce many.

The present inventor also recognized that the NIVD signals may includean undesirable amount of noise. To account for this noise, the presentinvention contemplates rejecting the NIVD signal for a particular breathbased on the signal to noise ratio (SNR) for the NIVD signal for thatbreath.

FIG. 21 illustrates a technique by which the present inventiondetermines a patient's pulsus paradoxis, which is the change in theenvelope of the pulse pressure (Cardiac Δd) throughout one breath. Thisfigures shows a raw NIVD (Δd) signal 215 produced by the extra-thoracicmonitoring system of the present invention over a plurality ofbreathings. As shown by arrows A and B, the percent change in the pulsepressure (Cardiac Δd) over one respiratory cycle can be determined. Forexample, this figure shows an approximately 32% change in pulse pressureover one breath. It can be appreciated that any technique forcalculating the change in pulse pressure, difference between the lengthof arrows A and B, can be used in the system of the present invention.

VIII. OTHER FEATURES OF THE EXTRA-THORACIC MONITORING SYSTEM

A. Other Sources for the Plethysmograph Signal

In the embodiment described above, the plethysmograph signal is obtainedoptically via a photodetector so that the plethysmography signal is aphotoplethysmography signal. It is to be understood that the presentinvention contemplates that other sensors can be used to monitor thechanges in the patient's circulatory system due to pressure or volumechanges in that system. For example, a blood pressure cuff that isdeflated is capable of detecting changes in vessel distention bydetecting the volume change in the vessel bed encompassed by the cuff.The pressure in the circulatory system can also be monitored directly byuse of an invasive pressure sensor, such as an arterial line, disposedin the patient. In short, any sensor that is capable of monitoring aphysiological characteristic of a patient associated with pressurechanges in such a patient's circulatory system is suitable to provideplethysmography signal 60 used by the processing system of the presentinvention.

B. Fractional Concentration of Inspired Oxygen

The present invention contemplates that the extra-thoracic monitoringsystem can include other functionalities and features. An example of afeature that can be added to the system is the capability to measure thefractional concentration of oxygen inhaled by a patient (FO₂) and thefractional concentration of oxygen inhaled by a patient over one breath(FIO₂). This measurement technique can be used alone or in conjunctionwith the cardio-pulmonary monitoring system discussed above.

As shown in FIG. 22, the FO₂ monitoring system 218 includes a patientcircuit 220 adapted to communicate a flow of breathing gas to an airwayof a patient. A first flow sensor 222 associated with the patientcircuit quantitatively measures a flow of gas (Q_(T)) inhaled, exhaled,or inhaled and exhaled by a patient. The FO₂ monitoring system alsoincludes an oxygen conduit 224 adapted to be coupled to an oxygen source226 and to the patient circuit to communicate oxygen from the oxygensource to such a patient. A second flow sensor 228 associated with theoxygen conduit quantitatively measures a flow of the oxygen (Q_(O2)) inthe oxygen conduit. In an exemplary embodiment, the pressure sensorsthat are used in flow sensors 222 and 228 are provided on pressure card66 of FIG. 3. A processing system (not shown) receives the signals fromthe flow sensors and determines the FO₂ based on the output of the firstflow sensor and the second flow sensor.

In one embodiment of the present invention, FO₂ at any given time iscalculated by the processor as follows:

${FO}_{2} = {\frac{{1.0\left( Q_{O\; 2} \right)} + {0.21\left( {Q_{T} - Q_{O\; 2}} \right)}}{Q_{T}} \cdot}$

This assumes that 100% oxygen is being delivered to the patient. If theoxygen concentration is less than 100%, the multiplier on Q_(O2) isadjusted to that concentration.

The processor calculates a fractional concentration of oxygen inhaled bya patient over one breath cycle (FIO₂) as follows:

${{FIO}_{2} = \frac{\int_{t_{1}}^{t_{2}}{\left( {FO}_{2} \right)\ {t}}}{t_{2} - t_{1}}},$

where t₁ corresponds to a time at a start of an inhalation phase of abreath cycle, and t₂ corresponds to a time at an end of the inhalationphase.

The present invention also contemplates determining FO₂ as follows:

${{FO}_{2} = \frac{{VO}_{2}}{V_{T}}},$

where VO₂ is the volume of oxygen delivered to the patient and isdetermined based on an output or the first and the second sensors asfollows:

VO₂ = ∫_(t₁)^(t₂)(Q_(O 2)) t + ∫_(t₁)^(t₂)(0.21(Q_(T) − Q_(O 2))) t,

where t₁ corresponds to a time at a start of an inhalation phase of abreath cycle, and t₂ corresponds to a time at an end of the inhalationphase, and where V_(T) is the volume of gas delivered to the patient andis determined based on an output of first sensor 222 as follows:

V_(T) = ∫_(t₁)^(t₂)Q_(T) t.

The processor calculates a fractional concentration of oxygen inhaled bya patient over one breath cycle (FIO₂) as discussed above.

FIG. 23 illustrates an FO₂ monitoring system 230 that is similar in manyrespects to the monitoring system shown in FIG. 22. The primarydifferences between monitoring system 230 and monitoring system 218 liesin the technique used to measure the flow of gas (Q_(T)) inhaled,exhaled, or inhaled and exhaled by the patient. In place of thepneumotach 222 shown in FIG. 22, monitoring system 230 in FIG. 23 uses aflow sensing system 232 disclosed in U.S. Pat. Nos. 6,544,192;6,342,040; and 6,017,315 all to Starr et al., the contents of each ofwhich are incorporated herein by reference, to measure the flow of gasinhaled and exhaled by the patient. In the embodiment, flow sensingsystem 232 measures the inhaled and exhaled air flow Q_(air) rather thanQ_(T). Thus calculating FO₂ is rewritten as:

${FO}_{2} = {\frac{{1.0\left( Q_{O\; 2} \right)} + {0.21\left( Q_{air} \right)}}{Q_{T}} \cdot}$

The FIO₂ is calculated as discussed above based on the measured FO₂.

C. Shunt Index Active Nomogram

Another example of a feature that can be added to the extra-thoracicmonitoring system is a system for displaying a nomograph that is used toestimate the percentage of a patient's shunt, also referred to as theshunt index. This display and estimation technique is used with the FIO₂measurement discussed above and the SpO₂ measurement that is obtainedfrom the photoplethysmography signal.

FIG. 24 illustrates an exemplary display 300 that can be shown, forexample, on display area 44 of monitoring system 30. A displaycontroller, such as the processor or the main computer shown in FIG. 1or 3, controls the display such that the display shows, in a first fieldon the display area, a nomogram illustrating a relationship between themeasured SpO₂, the FIO₂, and an estimated shunt. More specifically, thenomogram shows the SpO₂ on a first axis, the FIO₂ on a second axis, anda plurality of curves 302 a-302 e such that each curve corresponds to acommon estimated shunt percentage.

This is an “active” nomogram in that the display controller causes anindicator 304 to be displayed on the nomogram at a location defined bycoordinates corresponding to a current value of the SpO₂ and the FIO₂.That is, once the SpO₂ and the FIO₂ are determined, indicator 304 isplaced at the coordinates corresponding to these SpO₂ and the FIO₂values. This enables the user to quickly visualize which shunt indexcurve the indicator is close to, thereby providing the user with anestimation of the patient's shunt. The present invention furthercontemplates that the position of the indicator on the nomogram iscontinuously updated each time a new value for the SpO₂ or the FIO₂ isdetermined. As a further feature, the processor can calculate theestimated shunt based on the SpO₂ and the FIO₂ measurement, and thecalculated estimated shunt can be displayed as a numerical value in asecond field on the display area.

The present invention also contemplates showing one or more pastindicators on the nomogram along with the current indicator. The pastindicator(s) is displayed in the nomogram at a location defined bycoordinates of prior values for the SpO₂ and the FIO₂. This enables theuser to see how the patient's condition, SpO₂, FIO₂, and shunt index haschanged over time.

D. Device Screen Shots

FIGS. 25-33 are screen shots of a display in a user interface for usewith the extra-thoracic monitoring system of the present invention.These screens are displayed, for example, on display area 44 ofmonitoring system 30.

FIG. 25 illustrates a set up screen 330 that includes a patient datafield 332, a monitored parameter display field 334, an oxygenconcentration setting field 336, and a data display selection field 338.A note field (not shown) can also be provided. An exit selector 342 isan active field that allows the user to exit the set up screen. Patientdata field 332 is also an active field that allows the user to set thecharacteristics of the patient, such as patient identification number,patient interface (type of interface being used by that patient, such asmask, cannula, or pneumotach), monitoring mode, age, gender, weight,height, and patient type.

Monitored parameter display field 334 displays the parameters monitoredby the extra-thoracic monitoring system. These parameters can be updatedcontinuously or only as desired. The parameters include the following:

-   -   (1) Inspiratory Tidal Volume—volume of gas inspired by a patient        during the inspiratory phase of the breathing cycle (measured        via a flow sensor);    -   (2) Peak Inspiratory Flow—peak flow during the inspiratory phase        of the breathing cycle (measured via a flow sensor);    -   (3) Peak Expiratory Flow—peak flow during the expiratory phase        of the breathing cycle (measured via a flow sensor);    -   (4) Minute Ventilation—volume of gas inspired by the patient        during one minute (measured via a flow sensor);    -   (5) Respiratory Rate—breathing rate or breathing frequency—how        many breaths per minute the patient takes (measured via a flow        sensor);    -   (6) I:E Ratio—duration of inspiration versus duration of        inspiration for a breath (measured via a flow sensor);    -   (7) RSBI—breathing frequency versus tidal volume (calculated        value);    -   (8) Supplemental Liter Flow—(Q_(O2)) flow rate supplement gas,        such as oxygen, being delivered to the patient in addition to        the primary gas flow (measured via an oxygen flow sensor);    -   (9) FIO₂ Delivered—fractional concentration of oxygen inhaled by        a patient over one breath (calculated as discussed above);    -   (10) Inhaled Slope—slope of rise portion of inspiratory flow (an        example of this is shown in FIG. 17);    -   (11) Peak Inspiratory Pressure—(PIP) measured by a pressure        sensor;    -   (12) PEEP—Positive End Expiratory Pressure (measured via a        pressure sensor).    -   (13) Dynamic Airway Compliance—calculated from pressure and        flow;    -   (14) SpO₂—Oxygen saturation (measured from pulse oximeter (IR)        signal);    -   (15) Pulse Rate—heart rate (measured from pulse oximeter (IR)        signal);    -   (16) Estimated Shunt—calculated as discussed above;    -   (17) Cardiac Delta D—NIVD_(Cardiac) calculated as discussed        above;    -   (18) Thoracic Delta D—NIVD_(Thoracic) calculated as discussed        above; and    -   (19) Expiratory Tidal Volume—volume of gas expired by a patient        during the expiratory phase of the breathing cycle (measured via        a flow sensor);

Oxygen concentration setting field 336 is used to set the concentrationof oxygen of the gas being delivered to the patient. This isaccomplished according to one embodiment of the present invention bymoving triangular pointers 344 a and 344 b to the oxygen concentrationsetting. The upper setting 344 a is used to set the main gas oxygenconcentration, and the lower gas setting 344 b is used to set thesupplemental gas oxygen concentration. For example, if the patient isbreathing air without any supplemental oxygen, upper setting 344 a isset to 0.21 and lower setting 344 b is set to 0.21. If the patient isthen given pure oxygen supplemental to the main flow of air, the lowersetting is moved to 1.00. These settings are used, for example, incalculating FIO₂.

Data display selection field 338 is used to allow the user to selectother screens for display. There are two types of data displays: realtime displays, which show data as it's continuously monitored andcalculated, and trend displays, which show monitored or calculated dataover a period of time, such as over eight hours, accumulated on abreath-by-breath basis. Examples of real time displays are shown inFIGS. 26-29 and 31. Examples of trend displays are shown in FIGS. 30, 32and 33. Displaying these other data display screens in the display areais accomplished according to one exemplary embodiment of the presentinvention by providing a real time data pull-down menu 346 that containseach real time data screen selection and a trended data pull-down menu348 that contains each trended data screen selection.

As noted above, FIGS. 26-33 illustrate various screens that are used todisplay the information gathered and/or calculated by the monitoringsystem of the present invention in real time and as trended dataaccumulated over a period of time. It should be noted that these otherscreens allow the monitored variables and the calculated variables to bedisplayed quantitatively and graphically. The graphical presentation ofthis information is believed to be more easily perceived and understoodby a user than a simple quantitative display. It also allows changes inthe information, i.e., data trends, to be visualized clearly.

E. Wavelength Selection

One embodiment of the present invention uses the following twofrequencies of light: Red having a wavelength of approximately 660 nm,and Infrared having a wavelength of approximately 940 nm. However, thepresent invention contemplates that a single wavelength, at theisobestic point of approximately 805 nm, can be used in place of thesetwo frequencies.

If a wavelength of light is chosen that is not affected by oxygensaturation, the calculation for Δd is simplified. As shown in FIG. 34,isosbestic point 340 occurs at a wavelength of light where theextinction coefficients for oxyhemoglobin ε_(HbO2) and reducedhemoglobin ε_(Hb) are equal. (ε_(HbO2)=ε_(Hb)). The isosbestic point isat the wavelength 805 nm. Because oxygen saturation (SpO₂) does notaffect this wavelength, equation (1) can be simplified as follows:

$\begin{matrix}{{{NIVD} = {{\Delta \; d} = \frac{- {\ln \left( \frac{I_{X}}{I_{H}} \right)}}{{ɛ_{THb}(\lambda)} \cdot 2.265}}},} & (2)\end{matrix}$

where ε_(THb) is the extinction coefficient at the isosbestic point. Itshould be noted that C_(TotHb) represents the total hemoglobinconcentration of 2.265 milli-mole per liter (mM/L), which is based on apatient with the normal amounts of dyshemoglobin.

F. Identifying Respiratory Disorders

The FFT signal, i.e., the output of FFT transform 152 in FIG. 9,contains information that is useful in determining whether the patientsuffers from a respiratory disorder, such as obstructive apneas andcentral apneas. FIG. 35 illustrates a frequency spectrum signal 400 thatcorresponds, in general, to frequency spectrum signal 158 in FIG. 12B,and which is output from zero-out component 192. FIG. 35 alsoillustrates a first threshold 402 and a second threshold 404 that areused to determine whether the patient suffers from a respiratorydisorder.

Frequency spectrum signal 400 shown in FIG. 35 includes a first peak 406that corresponds to the breath rate component of the frequency spectrumsignal and a second peak 408 that corresponds to the heart ratecomponent. First threshold 402 corresponds to a threshold that, if metor exceeded by first peak 406 indicates that the patient suffers from anobstructive apnea. In other words, if the amplitude of the first peak4-6 meets or exceeds the first threshold, the patient is deemed to besuffering an obstructive apnea. An obstructive apnea is characterized bya high work of breathing, i.e., a high magnitude in first peak 406, buta minimal flow through the airway due to the obstruction of the airway.The patient is deemed to be suffering from no apneas if the amplitude ofthe respiratory portion 406 of FFT signal 400 is between the thresholds402 and 404.

A central apnea is declared if the amplitude of respiratory portion 406of FFT signal 400 meets or crosses below second threshold 404. A centralapnea is characterized by little or no work of breathing, but a minimalflow through the airway due to the central apnea. In an exemplaryembodiment of the present invention, first and second thresholds 402 and404 are calculated based on the FFT of a patient's normal breathing. Forexample, the present invention contemplates setting the first and secondthresholds based on an average of the previous peaks of the FFTs.

IX. CONCLUSION

It can be appreciated that the present invention provides a system formonitoring changes in the intra-thoracic pressure of a patient due tothe patient's respiratory activity or cardiac function in real time andon a continuous basis. For example, pleural pressure changes due torespiratory effort are monitored based on changes in pressure in theextra-thoracic arterial circulatory system to allow the caregiver toestimate work of breathing. The patient's blood pressure can also bemonitored continuously and non-invasively.

Although the invention has been described in detail for the purpose ofillustration based on what is currently considered to be the mostpractical and preferred embodiments, it is to be understood that suchdetail is solely for that purpose and that the invention is not limitedto the disclosed embodiments, but, on the contrary, is intended to covermodifications and equivalent arrangements that are within the spirit andscope of the appended claims.

1. A non-invasive cardiac monitoring system comprising: a photoemitteradapted to transmit light through a portion of a patient'sextra-thoracic arterial circulation; a photodetector adapted to receivelight after having been transmitted through such a patient and foroutputting a first signal based on the received light; processing meansfor producing a cardiac pressure signal as a measure of such a patient'scardiac function by isolating cardiac related pressure variations in thefirst signal.
 2. The system according to claim 1, wherein the processingmeans determines the cardiac pressure signal substantially continuously.3. The system according to claim 1, further comprising an outputtingmeans for presenting an indication of the cardiac pressure signal in ahuman perceivable format.
 4. The system according to claim 1, whereinthe processing means comprises: frequency analyzing means fordetermining frequency components of the first signal; heart ratefrequency component identifying means for identifying a heart ratefrequency component (f_(HR)) from the frequency components of the firstsignal; and dynamic filtering means for filtering the first signal basedon the heart rate frequency component f_(HR) so as to isolate the heartrate frequency component f_(HR) from the first signal to produce thecardiac pressure signal.
 5. The system according to claim 4, wherein thefrequency analyzing means is a means for performing a Fourier transformthat generates the frequency components of the first signal.
 6. Thesystem according to claim 4, wherein the processing means furthercomprises offset removing means for removing any offset in the frequencycomponents of the first signal.
 7. The system according to claim 4,further comprising heart rate monitoring means for determining a heartrate of such a patient, and wherein the heart rate frequency componentidentifying means selects the heart rate frequency component f_(HR) fromthe frequency components of the first signal based on the output of theheart rate monitoring means.
 8. The system according to claim 1, whereinthe processing means comprises: frequency analyzing means fordetermining frequency components of the first signal; respiratory ratefrequency component identifying means for identifying a respiratory ratefrequency component f_(RR) from the frequency components of the firstsignal; and dynamic filtering means for filtering the physiologic signalbased on the respiratory rate frequency component f_(RR) so as to removethe respiratory rate frequency component f_(RR) from the physiologicsignal, thereby isolating a heart rate frequency component f_(HR) of thefirst signal to produce the cardiac pressure signal.
 9. The systemaccording to claim 8, further comprising respiratory rate monitoringmeans for determining a respiratory rate of such a patient, and whereinthe respiratory rate frequency component identifying means selects therespiratory rate frequency component f_(RR) from the frequencycomponents of the first signal based on the output of the respiratorymonitoring means.
 10. The system according to claim 1, wherein theprocessing means comprises: frequency analyzing means for determiningfrequency components of the first signal, including a respiratory ratefrequency component f_(RR) and a heart rate frequency component f_(HR);respiratory rate frequency component f_(RR) identifying means foridentifying the respiratory rate frequency component f_(RR) from thefrequency components of the first signal; heart rate frequency componentf_(HR) identifying means for identifying the heart rate frequencycomponent f_(HR) from the frequency components of the first signal; anddynamic filtering means for filtering the first signal based on therespiratory rate frequency component f_(RR) and the heart rate frequencycomponent f_(HR) so as to isolate the heart rate frequency componentf_(RR) from the first signal to produce the cardiac pressure signal. 11.The system according to claim 10, wherein the dynamic filtering means(1) sets a cutoff frequency (f_(cutoff)) as f_(HR)−f_(smear) responsiveto respiratory rate frequency component f_(RR) being less than the heartrate frequency component f_(HR), (2) sets the cutoff frequency(f_(cutoff)) as f_(HR)+f_(smear) responsive to respiratory ratefrequency component f_(RR) being greater than the heart rate frequencycomponent f_(HR), where f_(smear) is a predetermined thresholdfrequency, and (3) high pass filters the first signal at the cutofffrequency f_(cutoff).
 12. The system according to claim 10, furthercomprising: respiratory rate monitoring means for determining arespiratory rate of such a patient, and wherein the respiratory ratefrequency component identifying means selects the respiratory ratefrequency component f_(RR) from the frequency components of the firstsignal based on the output of the respiratory rate monitoring means; andheart rate monitoring means for determining a heart rate of such apatient, and wherein the heart rate frequency component identifyingmeans selects the heart rate frequency component f_(HR) from thefrequency components of the first signal based on the output of theheart monitoring means.
 13. The system according to claim 1, wherein theprocessing means identifies an occurrence of at least one of pulsusparadoxis, pulsus alternans, pulsus bisferiens, dicrotic pulse,anacrotic pulse, “waterhammer” pulse or a normal pulse based on thecardiac pressure signal.
 14. A non-invasive cardiac monitoring methodcomprising: passing light through a portion of a patient'sextra-thoracic arterial circulation; receiving light after having beenpassed through such a patient; outputting a first signal based on thereceived light; and producing a cardiac pressure signal as a measure ofsuch a patient's cardiac function by isolating cardiac related pressurevariations in the first signal.
 15. The method according to claim 14,wherein the cardiac pressure signal is determined substantiallycontinuously.
 16. The method according to claim 14, further comprisingoutputting the cardiac pressure signal in a human perceivable format.17. The method according to claim 14, wherein producing a thoracicpressure signal comprises: determining frequency components of the firstsignal; identifying a heart rate frequency component (f_(HR)) from thefrequency components of the first signal; and filtering the first signalbased on the heart rate frequency component f_(HR) so as to isolate theheart rate frequency component f_(HR) from the first signal to producethe cardiac pressure signal.
 18. The method according to claim 17,wherein determining frequency components of the first signal includesanalyzing the first signal with a Fourier transform that generates thefrequency components of the first signal.
 19. The method according toclaim 17, wherein producing a cardiac pressure signal further comprisesremoving any offset in the frequency components of the first signal. 20.The method according to claim 17, further comprising determining a heartrate of such a patient via a heart rate monitor, and wherein identifyingthe heart rate frequency component f_(HR) from the frequency componentsof the first signal is accomplished based on the output of the heartrate monitor.
 21. The method according to claim 14, wherein producing athoracic pressure signal comprises: determining frequency components ofthe first signal; identifying a respiratory rate frequency componentf_(RR) from the frequency components of the first signal; and filteringthe physiologic signal based on the respiratory rate frequency componentf_(RR) so as to remove the respiratory rate frequency component f_(HR)from the physiologic signal, thereby isolating a heart rate frequencycomponent f_(HR) of the first signal to produce the thoracic pressuresignal.
 22. The method according to claim 21, further comprisingdetermining a respiratory rate of such a patient via a respiratory ratemonitor, and wherein identifying the respiratory rate frequencycomponent f_(RR) from the frequency components of the first signal isaccomplished based on the output of the respiratory monitor.
 23. Themethod according to claim 14, wherein producing a thoracic pressuresignal comprises: determining frequency components of the first signal,including a respiratory rate frequency component f_(RR) and a heart ratefrequency component f_(HR); identifying a respiratory rate frequencycomponent f_(RR) from the frequency components of the first signal;identifying a heart rate frequency component f_(HR) from the frequencycomponents of the first signal; and filtering the first signal based onthe respiratory rate frequency component f_(RR) and the heart ratefrequency component f_(HR) so as to isolate the heart rate frequencycomponent f_(RR) from the first signal to produce the cardiac pressuresignal.
 24. The method according to claim 23, wherein filtering thefirst signal includes: (1) setting a cutoff frequency (f_(cutoff)) asf_(HR)−f_(smear) responsive to the respiratory rate frequency componentf_(RR) being less than the heart rate frequency component f_(HR), (2)setting the cutoff frequency (f_(cutoff)) as f_(HR)+f_(smear) responsiveto the rate frequency component f_(RR) being greater than the heart ratefrequency component f_(HR), where f_(smear) is a predetermined thresholdfrequency, and (3) high pass filtering the first signal at the cutofffrequency f_(cutoff).
 25. The method according to claim 23, furthercomprising: determining a respiratory rate of such a patient via arespiratory rate monitor, and wherein identifying the respiratory ratefrequency component f_(RR) from the frequency components of the firstsignal is accomplished based on the output of the respiratory ratemonitor; and determining a heart rate of such a patient based on a heartrate monitor, and wherein identifying the heart rate frequency componentf_(HR) from the frequency components of the first signal is accomplishedbased on the output of the heart rate monitor.
 26. The system accordingto claim 14, further comprising identifying an occurrence of at leastone of pulsus paradoxis, pulsus alternans, pulsus bisferiens, dicroticpulse, anacrotic pulse, “waterhammer” pulse or a normal pulse based onthe cardiac pressure signal.
 27. A cardio-pulmonary monitoring systemcomprising: (a) sensing means for detecting a physiologicalcharacteristic of a patient associated with pressure changes in such apatient's extra-thoracic arterial circulation and for outputting a firstsignal indicative of such pressure changes; (b) processing means forproducing a cardiac pressure signal as a measure of such a patient'scardiac function by isolating cardiac related pressure variations in thefirst signal, wherein the processing means comprises: (1) frequencyanalyzing means for determining frequency components of the firstsignal, (2) heart rate frequency component identifying means foridentifying a heart rate frequency component (f_(HR)) from the frequencycomponents of the first signal, and (3) dynamic filtering means forfiltering the first signal based on the heart rate frequency componentf_(HR) so as to isolate the heart rate frequency component f_(HR) fromthe first signal to produce the cardiac pressure signal.
 28. The systemaccording to claim 27, further comprising heart rate monitoring meansfor determining a heart rate of such a patient, and wherein the heartrate frequency component identifying means selects the heart ratefrequency component f_(HR) from the frequency components of the firstsignal based on the output of the heart monitoring means.
 29. The systemaccording to claim 27, wherein the heart rate monitoring means comprisesa pulse oximeter.
 30. The system according to claim 27, wherein theprocessing means determines the cardiac pressure signal substantiallycontinuously.
 31. The system according to claim 27, further comprisingan outputting means for presenting an indication of the cardiac pressuresignal in a human perceivable format.
 32. A cardio-pulmonary monitoringsystem comprising: (a) sensing means for detecting a physiologicalcharacteristic of a patient associated with pressure changes in such apatient's extra-thoracic arterial circulation and for outputting a firstsignal indicative of such pressure changes; (b) processing means forproducing a cardiac pressure signal as a measure of such a patient'scardiac function by isolating cardiac related pressure variations in thefirst signal, wherein the processing means comprises: (1) frequencyanalyzing means for determining frequency components of the firstsignal, (2) respiratory rate frequency component identifying means foridentifying a respiratory rate frequency component f_(RR) from thefrequency components of the first signal, and (3) dynamic filteringmeans for filtering the physiologic signal based on the respiratory ratefrequency component f_(RR) so as to remove the respiratory ratefrequency component f_(RR) from the physiologic signal, therebyisolating a heart rate frequency component f_(HR) of the first signal toproduce the cardiac pressure signal.
 33. The system according to claim32, further comprising respiratory rate monitoring means for determininga respiratory rate of such a patient, and wherein the respiratory ratefrequency component identifying means selects the respiratory ratefrequency component f_(RR) from the frequency components of the firstsignal based on the output of the respiratory monitoring means.
 34. Thesystem according to claim 32, wherein the respiratory rate monitoringmeans comprises a flow sensor adapted to detect a flow of gas two orfrom such a patient.